Download Independent effects of preload, afterload, and contractility on left

Independent effects of preload, afterload, and
contractility on left ventricular torsion
SHENG-JING DONG, PAUL S. HEES, WEN-MEI HUANG, SAM A. BUFFER, JR.,
JAMES L. WEISS, AND EDWARD P. SHAPIRO
Division of Cardiology, Johns Hopkins University School of Medicine,
Johns Hopkins Bayview Medical Center, Baltimore, Maryland 21224
Dong, Sheng-Jing, Paul S. Hees, Wen-Mei Huang,
Sam A. Buffer, Jr., James L. Weiss, and Edward P.
Shapiro. Independent effects of preload, afterload, and contractility on left ventricular torsion. Am. J. Physiol. 277
(Heart Circ. Physiol. 46): H1053–H1060, 1999.—Shortening
of oblique left ventricular (LV) fibers results in torsion. A
unique relationship between volume and torsion is therefore
expected, and the effects of load and contractility on torsion
should be predictable. However, volume-independent behavior of torsion has been observed, and the effects of load on this
deformation remain controversial. We used magnetic resonance imaging (MRI) with tagging to study the relationships
between load and contractility, and torsion. In ten isolated,
blood-perfused canine hearts, ejection was controlled by a
servopump: end-diastolic volume (EDV) was controlled by
manipulating preload parameters and end-systolic volume
(ESV) by manipulating afterload using a three-element windkessel model. MRI was obtained at baseline, two levels of
preload alteration, two levels of afterload alteration, and
dobutamine infusion. An increase in EDV resulted in an
increase in torsion at constant ESV (preload effect), whereas
an increase in ESV resulted in a decrease in torsion at
constant EDV (afterload effect). Dobutamine infusion increased torsion in association with an increase in LV peaksystolic pressure (PSP), even at identical EDV and ESV.
Multiple regression showed correlation of torsion with preload (EDV), afterload (ESV), and contractility (PSP; r ⫽ 0.67).
Furthermore, there was a close linear relationship between
torsion and stroke volume (SV) and ejection fraction (EF)
during load alteration, but torsion during dobutamine infusion was greater than expected for the extent of ejection.
Preload and afterload influence torsion through their effects
on SV and EF, and there is an additional direct inotropic effect
on torsion that is independent of changes in volume but
rather is force dependent. There is therefore potential for the
torsion-volume relation to provide a load-independent measure of contractility that could be measured noninvasively.
left ventricle; twist; magnetic resonance imaging
CARDIAC TORSION is the relative rotation of the left
ventricular (LV) apex with respect to the base, about LV
long axis. During the past decade, this deformation has
been widely studied and quantified in both animal and
human models with various methods (2–7, 10, 11, 13,
14, 20, 21). Because torsion results from shortening of
obliquely oriented myocardial fibers (3, 15), the presence of a unique relation between cavity volume and
the extent of torsion is expected and has been reported
The costs of publication of this article were defrayed in part by the
payment of page charges. The article must therefore be hereby
marked ‘‘advertisement’’ in accordance with 18 U.S.C. Section 1734
solely to indicate this fact.
(2). The effects of load and contractility on torsion
should therefore be predictable. However, these remain
controversial. Studies in transplanted human hearts
(13, 21) suggest that torsion, as measured by radiopaque markers, is not affected by pressure or volume
loading, whereas experiments in dog models report
that torsion, as measured by markers (2) or an optical
device (10, 11), is primarily a function of volume at end
diastole and end systole and is preload and afterload
dependent. Furthermore, although it has been well
demonstrated that torsion is very sensitive to changes
in contractility (6, 11, 13, 14, 21) (positive inotropic
interventions increase torsion, whereas negative inotropic interventions decrease torsion), it has been difficult
to distinguish whether this inotropic effect is mediated
through (11) or is independent of (13) changes in
volume.
A volume-independent component of torsion has been
clearly observed; substantial recoil of torsion is known
to occur between the time of aortic valve closure and
mitral valve opening, when cavity volume is fixed (10,
11, 23), and systolic torsion has been demonstrated in
isolated isovolumic beating hearts (20). Thus under
some circumstances, torsion may vary with force, as
well as with volume. Because inotropic stimulation
enhances pressure (or force) generation at any given
end-diastolic volume (EDV) and end-systolic volume
(ESV; 17, 18), it may therefore have a direct effect on
torsion that is not mediated through changes in volume
but through changes in force.
To examine further the individual effects of preload,
afterload, and contractility on torsion, we used magnetic resonance imaging (MRI) with tissue tagging (29),
which permits the accurate measurement of torsion, to
study an isolated, blood-perfused, ejecting heart model
that allows independent control of these factors. We
hypothesized that 1) preload and afterload, through
their effect on volume, affect torsion and 2) there is a
further direct inotropic effect on torsion that is independent of changes in volume but rather is force dependent.
METHODS
Animal Preparation
Ten pairs of mongrel dogs were anesthetized with pentobarbital sodium (25–35 mg/kg iv). The smaller dog of the pair
served as the heart donor (mean wt 21.2 ⫾ 1.5 kg) and the
larger one as the support dog (mean wt 30.1 ⫾ 1.5 kg).
Intravenous heparin (10,000 U), hydrocortisone (250 mg),
and indomethacin (25 mg) were administered to the support
dog, which was mechanically ventilated (model 613; Harvard
Apparatus, South Natick, MA).
0363-6135/99 $5.00 Copyright r 1999 the American Physiological Society
H1053
H1054
DETERMINANTS OF LEFT VENTRICULAR TORSION
The procedures to isolate and support the heart were
similar to those previously described (20, 25). Briefly, the
femoral arteries and veins of the support dog were cannulated
and connected to a perfusion system, which was primed with
1 liter of 50% dextran in saline mixed with the blood of the
support dog. The arterial pressure of the support dog was
continuously monitored with a fluid-filled catheter placed in
the aorta, and serial arterial blood gases were taken and any
deficiencies corrected accordingly.
The chest of the donor dog was opened under artificial
ventilation. The left subclavian artery was cannulated with
the arterial perfusion line and the brachiocephalic artery was
cannulated to monitor coronary perfusion pressure. Filtered,
warmed (37°C), and oxygenated arterial blood from the
femoral arteries of the support dog was supplied to the
isolated heart through a pump (model 7518–00; Cole Parmer
Instrument, Chicago, IL) under pressure, which was monitored by a fluid-filled catheter and maintained at 80 mmHg by
adjusting the speed of the pump. Once the heart was crossperfused, the azygos vein, superior and inferior venae cavae,
descending aorta, and lung hila were ligated, and the heart
was removed from the thoracic cavity of the donor dog. The
coronary venous blood and any accumulated blood in the left
ventricle were drained through vents placed in the respective
right and LV apex into a receptacle, and via a second pump it
was returned to the femoral veins of the support dog. The left
atrium was opened, the chordae tendineae were detached
from the mitral valve leaflets, and a plastic ring that holds the
isolated heart onto the servopump apparatus was sutured
into the mitral valve ring. Electrocardiogram (ECG) leads
and atrial pacing leads were sutured onto the heart. The
isolated heart was then fixed to the servopump horizontally
with a water-filled latex balloon inside the LV cavity. LV
pressure was monitored with a fluid-filled catheter placed
within the balloon.
Load Control Via Servopump System
Load control was carried out by a servopump system
(Vivitro Systems, Victoria, Canada), which includes a linear
motor, an MRI compatible cylinder pump, a 100-cm long
plastic hollow shaft connecting the motor and piston of the
cylinder, and a control box. A latex balloon was secured to a
plastic tube connected to the fluid port of the cylinder. The
cylinder, plastic tube, and balloon were filled with tap water.
This servopump system allowed the manipulation of preload
via a computer-simulated preload circuit consisting of a filling
resistance and pressure source (25), affecting EDV, and the
manipulation of afterload via a computer-simulated afterload
analog consisting of a characteristic impedance, peripheral
arterial resistance, and arterial compliance (3-element windkessel), affecting ESV (12, 26, 27). During the filling phase,
the instantaneous LV pressure is compared with the desired
left atrial (LA) pressure. The difference between LA and LV
pressure is divided by the mitral valve resistance to obtain
flow. As long as LV pressure exceeds LA pressure, the flow is
zero. If LV pressure is lower than LA pressure, the flow is
integrated and a filling volume is obtained. During ejection,
LV pressure is compared with simulated aortic pressure. The
difference between LV and aortic pressure is divided by the
characteristic impedance to obtain flow. This flow is integrated to obtain a volume that is allowed to be ejected from
the LV through movement of the pump’s piston. The design
and implementation details of this servopump system have
been reported previously (20, 25).
Experimental Protocol
Protocol 1: load alteration. With this experimental preparation, the computer simulation of load may be altered, causing
secondary changes in EDV and ESV. In eight hearts, the
loading conditions were changed under a constant inotropic
status (Fig. 1). We first changed the preload (preload down
and preload up) so that EDV during intervention of preload
down and ESV during intervention of preload up were
arbitrarily maintained at a value that was midway between
baseline EDV and ESV. We then changed the afterload, such
that the ESV during intervention of afterload down was
matched to that during intervention of preload down, and the
ESV during intervention of afterload up was matched to that
during intervention of preload up, whereas EDV was maintained the same as that of baseline. This combination of
volumes was designed to allow assessment of preload effects
by comparison of runs with different EDV but identical ESV
(i.e., intervention 2 was compared with intervention 4, and
intervention 3 to intervention 5). It also allowed determination of afterload effects by comparison of runs with identical
EDV but different ESV (i.e., interventions 1, 4, and 5).
Protocol 2: contractility alteration. In eight hearts, contractility was increased by dobutamine infusion, whereas EDV
and ESV were maintained identical to those of baseline by
adjusting the preload and afterload (Fig. 1). Dobutamine was
infused through the support dog, and its dosage was adjusted
to increase LV peak-systolic pressure (PSP) by at least 30%.
MRI Acquisition and Analysis
Image acquisition. MRI was performed on a Resonex 0.38
tesla iron-core resistive magnet (Resonex, RX4000; Sunnyvale, CA) with a receiver-transmitter radio frequency coil
that provides enhanced image resolution and signal-to-noise
ratios (24). Image acquisition includes a gradient echo sequence modified to include tissue tagging, with temporal
resolution of 16 ms, time to echo of 10 ms, time to repeat of
two R-R intervals, 128 phase, and 256 frequency encoding
steps, 2–4 averages of excitation, and 20–25 cm field of view.
Two (basal and apical) short-axis images, 3.0 cm apart, with
eight radial tags, perpendicular to the image planes, were
acquired. Tagging and imaging were triggered by the R wave
of the ECG; tags were placed at end diastole and 25 images
were acquired thereafter. The image slice thickness was 1.0
cm and tag thickness 0.3 cm. Heart rate was kept constant by
Fig. 1. Schematic illustration of volume control during each intervention. Note that 1) end-systolic volume (ESV) was held constant in
interventions 2 and 4 and interventions 3 and 5, whereas the
end-diastolic volume (EDV) was changed; 2) EDV was held constant
in interventions 1, 4, and 5, whereas ESV was varied; and 3) all
volumes were held constant in interventions 6 and 7.
DETERMINANTS OF LEFT VENTRICULAR TORSION
atrial pacing. Figure 2 shows the end-diastolic and endsystolic basal and apical images.
Image analysis. The files containing the digitized MRI data
were transferred through a network to an IBM-compatible
personal computer (Gateway 2000 P5–90; North Sioux City,
SD), where image processing and data analysis were performed by use of a special software package (Cardiology
Image Processing System, Johns Hopkins University).
The tag-endocardium intersection points on both basal and
apical short-axis images were identified manually and digitized (8 points per slice). It should be noted that torsion of the
endocardial surface, which is larger than that of the epicardial surface, was used in this study. An Akima smoothing
algorithm (1) was used to interpolate between the intersection points and to smooth the contour.
Torsion was calculated as an angle between the tag line of
apical slice and the projection of the basal tag line on the
apical slice (Fig. 3), using a technique that is similar to that
previously reported (7). Briefly, the centroid of the intersection points on each image was calculated, and then the slopes
of the lines connecting each tag-endocardial intersection
point and the corresponding centroid were calculated and
expressed as an angle. To subtract out the effect of rigid body
rotation and whole heart translation, the image of the basal
plane at each time point was used as a reference. The
difference in the slope of the line connecting an intersection
point on the basal plane with its centroid and the line
connecting the corresponding intersection point to its centroid on the apical plane was calculated, and its arctangent
was called the torsion angle (␪). This angle is the difference, at
each time point, between the position of a basal intersection
point and the corresponding apical intersection point, expressed as an angle of rotation. Because there was no
difference in the relative positions of these two points at end
diastole, the tags being inserted as a plane in both slices
simultaneously, this angle represents the rotation of one
point with respect to the other between the time of tag
insertion and the time of measurement. Increasing angle
(positive sign) indicates counterclockwise rotation of the apex
when viewed looking from the apex toward the base (Fig. 3).
Mean torsion angles were then calculated as the average of
eight values for the eight tag-endocardial intersection points.
Assuming constant arc length change, we related the
torsion angle (␪) to the short-axis radius of the LV cavity (7). It
is necessary to take this into account when comparing torsion
at different cavity volumes, as occurs when loads are varied.
Circumferential-longitudinal shear (CL shear) is the difference between the systolic position of an apical tag point and
H1055
Fig. 3. Schematic illustration of calculation of torsion (␪) and circumferential-longitudinal shear (␥) for one tag point; r is short-axis
radius of cavity and h is distance between basal and apical image
planes.
the equivalent tag point on the basal slice, expressed as an
angle (␥, Fig. 3) of which the vertex is the tag point on the
basal slice, and calculated as CL shear equals tan⫺1[2 · r · sin(␪/
2)/h] (7), where r is the radius calculated from the area (A)
within the endocardial contour, as r ⫽ (A/␲)1/2, and h is the
distance between basal and apical slices.
Statistical Analysis
Results are expressed as means ⫾ SD. Repeated-measures
analysis of variance with the Student-Newman-Keuls correction was used to determine the statistical significance of the
interventions. Multiple linear regression was used to determine the effect of EDV, ESV, and dobutamine infusion on
torsion, and simple linear regression was used to correlate
torsion and stroke volume (SV) and ejection fraction (EF). A P
value of ⬍ 0.05 was considered to be statistically significant.
RESULTS
Fig. 2. End-diastolic (A) and end-systolic (B) basal and apical
short-axis images. Note that 8 radially oriented tissue tags intersect
left ventricle, producing 8 segments.
Preload alteration changed significantly both EDV
and ESV. Compared with baseline, there was no significant difference in EF, although SV decreased during
preload down. Afterload alteration changed only ESV,
resulting in changes in SV and EF (Table 1). During
load alteration, the heart rate was paced at a constant
rate (142 ⫾ 3 beats/min). Dobutamine infusion increased PSP, even though EDV and ESV, and thus the
SV and EF, were maintained identical to those at
baseline condition (Table 1). Because of the chronotropic effect of dobutamine, the heart rate was slightly
higher than the baseline pacing rate in four of eight
dogs and therefore was paced at this higher rate, but
there was no significant difference between baseline
and dobutamine infusion (152 ⫾ 12 vs. 163 ⫾ 13
beats/min).
H1056
DETERMINANTS OF LEFT VENTRICULAR TORSION
Table 1. Left ventricular pressure and volume response to intervention
EDP,
mmHg
PSP,
mmHg
EDV,
ml
ESV,
ml
SV,
ml
EF,
%
Baseline
PL down
PL up
AL down
AL up
5.5 ⫾ 2.3
3.1 ⫾ 1.4*
8.5 ⫾ 2.8*
5.3 ⫾ 1.9
5.6 ⫾ 2.0
104 ⫾ 13
89 ⫾ 10*
121 ⫾ 13*
84 ⫾ 11*
109 ⫾ 13
26.7 ⫾ 6.4
21.7 ⫾ 6.8*
32.4 ⫾ 5.2*
26.7 ⫾ 6.2
26.7 ⫾ 6.4
17.1 ⫾ 6.7
14.3 ⫾ 6.9*
21.9 ⫾ 6.5*
14.5 ⫾ 6.4*
21.5 ⫾ 7.1*
9.5 ⫾ 1.3
7.4 ⫾ 1.2*
10.5 ⫾ 2.7
12.2 ⫾ 1.7*
5.2 ⫾ 1.1*
38.5 ⫾ 14.0
38.1 ⫾ 16.2
33.8 ⫾ 12.4
48.4 ⫾ 15.5*
21.5 ⫾ 9.6*
Baseline
Dobutamine
4.9 ⫾ 2.5
5.9 ⫾ 1.1
98 ⫾ 17
160 ⫾ 25*
26.4 ⫾ 5.8
26.4 ⫾ 5.5
17.0 ⫾ 5.9
16.7 ⫾ 5.4
9.4 ⫾ 1.0
9.7 ⫾ 0.6
37.6 ⫾ 11.4
38.6 ⫾ 10.1
Values are means ⫾ SD; n ⫽ 8 dogs. EDP, end-diastolic pressure; PSP, peak systolic pressure; EDV, end-diastolic volume; ESV, end-systolic
volume; SV, stroke volume; EF, ejection fraction; PL, preload; AL, afterload. * P ⬍ 0.05 compared with baseline.
Effect of Preload on Torsion
At constant ESV, the torsion was greater at a higher
EDV (Fig. 4). At ESV of 14 ml, torsion was 9.4 ⫾ 3.4 and
12.2 ⫾ 4.0° (P ⬍ 0.01) at EDV of 21.7 ⫾ 6.7 and 26.7 ⫾
6.2 ml, respectively. At ESV of 22 ml, torsion was 6.4 ⫾
5.1 and 8.1 ⫾ 3.0° (P ⬍ 0.05) at EDV of 26.7 ⫾ 6.4 and
32.4 ⫾ 5.2 ml, respectively. Similar results were also
found in CL shear: 2.5 ⫾ 0.6 vs. 3.3 ⫾ 0.7° (P ⬍ 0.01),
and 1.9 ⫾ 1.0 vs. 2.6 ⫾ 0.8° (P ⬍ 0.01), respectively.
P ⬍ 0.05) (Fig. 6) and CL shear (3.4 ⫾ 1.1 vs. 2.5 ⫾ 0.7°,
P ⬍ 0.05) in association with an increase in LV PSP
(160 ⫾ 25 vs. 98 ⫾ 17 mmHg, P ⬍ 0.05, Table 1).
Multiple Regression Analysis
To determine further the effect of preload, afterload,
and contractility on torsion, we pooled data from all
hearts and performed a multiple linear regression
analysis using the model
Torsion ⫽ a0 ⫹ a1 · EDV ⫹ a2 · ESV ⫹ a3 · Di
Effect of Afterload on Torsion
At constant EDV (26.7 ml), the torsion decreased
from 12.2 ⫾ 4.0 to 10.7 ⫾ 3.9 to 6.4 ⫾ 5.1° (P ⬍ 0.001) as
the ESV increased from 14.5 ⫾ 6.4 to 17.1 ⫾ 6.7 to
21.5 ⫾ 7.1 ml, due to the increased afterload (Fig. 5).
Similar results were also observed in CL shear, which
decreased from 3.3 ⫾ 0.7 to 3.0 ⫾ 0.7 to 1.9 ⫾ 1.0° (P ⬍
0.001).
Inotropic Effect on Torsion
At identical EDV (26 ml) and ESV (17 ml), and
therefore identical SV and EF, dobutamine infusion
significantly increased torsion (13.3 ⫾ 4.3 vs. 8.7 ⫾ 1.5°,
Fig. 4. Effect of preload on torsion at constant ESV. At constant ESV,
torsion was greater at higher EDV.
or
CL shear ⫽ a0 ⫹ a1 · EDV ⫹ a2 · ESV ⫹ a3 · Di
where Di is a dummy variable to account for inotropic
status: Di equals 0 for baseline inotropic state and Di
equals 1 for dobutamine infusion. We found that there
was a significant correlation between torsion and EDV,
ESV, and inotropic status (multiple regression: r ⫽
0.66, P ⬍ 0.0001). As shown in Table 2, EDV was a
positive predictor (P ⬍ 0.05), i.e., the greater the EDV,
the greater the torsion, whereas ESV is a negative
predictor (P ⬍ 0.0001), i.e., the greater the ESV, the
smaller the torsion. Note that the absolute value of the
coefficient for EDV is about two-thirds as great as that
for ESV, indicating that the effect of EDV on the torsion
is about two-thirds as great as that of ESV. The
dobutamine infusion also has a significant positive
effect on torsion (P ⬍ 0.01). Similarly, significant correlation was also found between CL shear and EDV, ESV,
and dobutamine infusion (multiple regression: r ⫽ 0.67,
P ⬍ 0.0001; Table 2).
Fig. 5. Effect of afterload on torsion at constant EDV. At constant
EDV, torsion decreased as ESV increased.
DETERMINANTS OF LEFT VENTRICULAR TORSION
H1057
Fig. 6. Inotropic effect on torsion at constant volumes. At identical
EDV and ESV, dobutamine infusion significantly increased torsion.
We also performed similar multiple linear regression
analysis using LV PSP instead of coded dummy variable for inotropic status
Torsion ⫽ b0 ⫹ b1 · EDV ⫹ b2 · ESV ⫹ b3 · PSP
or
CL shear ⫽ b0 ⫹ b1 · EDV ⫹ b2 · ESV ⫹ b3 · PSP
Similar results were found and are presented in Table 2.
Relationship Between Torsion and SV and EF
A direct linear relationship between torsion and SV
and EF was detected from data of baseline and load
alterations (torsion ⫽ 3.5 ⫹ 0.62 · SV, r ⫽ 0.75, and
torsion ⫽ 1.37 ⫹ 0.22 · EF, r ⫽ 0.94, Fig. 7, open circles),
indicating that the greater the shortening, the greater
the torsion. However, data from dobutamine infusion
fell upward to and outside of the 95% confidence
interval of the relations (Fig. 7, filled circles), indicating
that the torsion was greater with dobutamine infusion
even at the same SV and EF. Similar relations were
found between CL shear and SV and EF (CL shear ⫽
1.04 ⫹ 0.18 · SV, r ⫽ 0.90 and CL shear ⫽ 0.78 ⫹
0.05 · EF, r ⫽ 0.93 during load alterations, and data
from dobutamine infusion were likewise shifted upward).
DISCUSSION
By using MRI tissue tagging and the isolated, bloodperfused, ejecting heart model, we studied the independent effects of preload, afterload, and contractility on
Table 2. Results of multiple regression analyses
Dependent
Variable
Torsion
CL shear
Torsion
CL shear
Independent
Variable
Regression
Coefficient
P
EDV
ESV
Dobutamine
EDV
ESV
Dobutamine
0.472 ⫾ 0.181
⫺0.731 ⫾ 0.173
3.429 ⫾ 1.235
0.118 ⫾ 0.040
⫺0.176 ⫾ 0.038
0.622 ⫾ 0.273
0.0117
⬍0.0001
0.0076
0.0048
⬍0.0001
0.0267
EDV
ESV
PSP
EDV
ESV
PSP
0.442 ⫾ 0.169
⫺0.761 ⫾ 0.161
0.059 ⫾ 0.015
0.114 ⫾ 0.039
⫺0.182 ⫾ 0.037
0.010 ⫾ 0.003
0.0119
⬍0.0001
0.0002
0.0053
⬍0.0001
0.0069
Values are means ⫾ SD; n ⫽ 8 dogs.
Fig. 7. Relationship between torsion and stroke volume (SV) and
ejection fraction (EF). There was significant correlation between
torsion and SV (A) and between torsion and EF (B) during load
alteration (s), whereas the data point during dobutamine infusion
was shifted upward, outside the 95% confidence interval (r).
mean LV endocardial torsion. We found that 1) preload
affects torsion: torsion was greater at higher EDV when
ESV was held constant (Fig. 4), 2) afterload affects
torsion: torsion decreased at higher ESV when EDV
was held constant (Fig. 5), 3) the effect of preload (as
assessed by EDV) on torsion is about two-thirds as
great as that of afterload (as assessed by ESV) (Table
2), and 4) contractility has a direct volume-independent
effect on torsion: dobutamine infusion increased torsion
even at identical EDV and ESV (Fig. 6). The independent effect of contractility on torsion was further demonstrated by multiple linear regression analysis.
Effect of Preload on Torsion
In contrast to the results of this study, it has been
previously reported that torsion is insensitive to preload. Hansen et al. (13, 14), using implanted midwall
radiopaque markers, did not find change in torsion
after volume loading in the transplanted human heart.
Similarly, Gibbons Kroeker et al. (10, 11), using an
optical apex rotation measuring device, did not observe
change in the magnitude of apical rotation either
during volume loading with saline or venae cavae
occlusion in the dog. However, in these studies of
preparations with intact circulation, volume expansion
and vena caval occlusion altered afterload as well as
preload in the same direction, i.e., volume load increases both preload and afterload (21), while vena
caval occlusion decreases both preload and afterload
(10, 11). Thus both EDV and ESV increase during
volume loading (13), whereas both EDV and ESV
decrease during vena caval occlusion (10, 11). These
H1058
DETERMINANTS OF LEFT VENTRICULAR TORSION
mixed effects account for the unchanged magnitude of
torsion in the intact ventricle. The current study demonstrates an important effect of pure preload alteration
on torsion.
Effect of Afterload on Torsion
In this study, we found a decrease in torsion in
response to increased afterload. This result is consistent with two previous dog studies in which preload
was kept constant; MacGowan et al. (20) increased
afterload by using an isolated heart model similar to
the present study, whereas Gibbons Kroeker et al. (11)
increased afterload by single beat aortic constriction in
dogs with intact circulation. However, in human transplanted heart studies (13, 21), no change in torsion was
observed with methoxamine-induced pressure loading.
This observation might result from the fact that a
methoxamine-induced increase in afterload could lead
to a secondary preload increase as suggested by the
increased EDV and end-diastolic pressure (21) and
unchanged EF (13, 21) in the patients studied.
The direct linear relationship between torsion and
SV and EF is in agreement with the concept that there
is a close correlation between torsion and volume,
independent of loading conditions, advanced by Arts et
al. (2, 3).
From the multiple linear regression analysis, we
found that the slope of torsion-EDV relation is about
two-thirds as great as that of torsion-ESV relation (the
absolute value of regression coefficient is 0.42 ⫾ 0.18 vs.
0.73 ⫾ 0.17, Table 2), indicating that the effect of
preload on torsion is about two-thirds as great as that
of afterload. This result is similar to that of an intact
dog study that showed that the mean slope of rotationcavity area relation at end diastole is about one-half as
great as that at end systole. (11) This effect is also seen
in studies of individual subjects where torsion-volume
loops are plotted through the cardiac cycle; during
systole the torsion-volume relation is steep, during
isovolumic relaxation there is sharp recoil without
volume change, and during diastole, the torsionvolume relation is gradual (21).
Effect of Contractility on Torsion
The overall effect of contractility on torsion has been
well demonstrated; positive inotropic interventions,
such as dobutamine infusion and paired pacing, greatly
increase torsion (6, 11, 13, 21), whereas negative inotropic interventions, such as ischemia and rejection reaction of the transplanted heart, markedly decrease
torsion (6, 14). In the intact circulation, changes in
contractility are accompanied by changes in SV and EF.
Therefore, it is difficult to distinguish whether this
inotropic effect is mediated through or is independent
of changes in volume. We observed here that dobutamine infusion increases torsion, in association with an
increase in PSP, even at identical EDV and ESV,
indicating that there is a direct inotropic effect on
torsion that is not mediated through changes in volume
but through changes in force. This effect was not
predicted by the modeling studies of Arts et al. (2, 3)
and was not observed by Gibbons Kroeker et al. (11)
during paired-pacing, a weaker inotropic stimulus.
The dependence of a global myocardial deformation
on the force of contraction, independent of changes in
volume, is unusual. However, the existence of a volumeindependent component of torsion has been well established previously by several investigators. For example, Beyar et al. (5), Rademakers et al. (23), and
Moon et al. (21) have shown that during isovolumic
relaxation, when cavity volume is fixed, there is a rapid
recoil of about 40% of the torsion that has accumulated
during systole. In addition, MacGowan et al. (20)
observed that considerable torsion persisted in the
isovolumic beating heart despite the absence of thickening, or circumferential or longitudinal shortening. Principal strain, to which torsional deformation contributes, also persisted in the isovolumic model.
The mechanism of this volume independence and
force dependence of torsion is uncertain. It appears that
when increased contractility due to ␤1 stimulation
cannot achieve greater cavity volume reduction, the
force along the oblique fiber direction is applied toward
increased intramyocardial deformation, that is, shear
in the CL direction. This interpretation is consistent
with the findings of Young et al. (28), who reported that
in patients with hypertrophic cardiomyopathy, minimal principal strain magnitude is preserved in association with an increased CL shear, despite decreases in
normal circumferential and longitudinal strains and
suggested that ‘‘the mechanical work is contributing to
wall shearing and not to cavity volume reduction.’’ This
shearing is likely to be a manifestation of the further
rearrangement of fiber sheets, which slide along ‘‘cleavage planes’’ within the myocardium during systole (19).
Our data suggest that this altered configuration can
occur entirely intramyocardially, independently of inward endocardial wall motion, and that it may contribute importantly to force development. It is possible that
such deformation may occur in association with greater
systolic compression of myocardium due to further
extrusion of blood from the intramyocardial vessels.
That is, myocardial volume may be reduced in systole
to a greater extent with inotropic stimulation, allowing
more wall deformation to occur. The absolute reduction
in myocardial volume during systole is small, less than
4% at baseline (16) , and has been difficult to detect by
MRI methods.
Another possible mechanism for altered wall deformation without an overall cavity volume change, might be
a change in LV cavity shape during inotropic stimulation. Altered shape might result in a redistribution of
myocardial mass, such that increased torsion could
occur. A shape analysis could not be performed in this
study since long-axis images were not obtained, due to
time constraints imposed by the fragile isolated heart
preparation. Olsen et al. (22) measured both short- and
long-axis diameters using implanted ultrasound crystals and reported that inotropic stimulation did not
alter the relationship between LV eccentricity and
volume; however, eccentricity does not fully describe LV
DETERMINANTS OF LEFT VENTRICULAR TORSION
H1059
shape. Future MRI study of torsion should include long
axis imaging and detailed shape analysis, so that the
interrelationships among shape, volume, and torsion
can be evaluated.
Finally, torsion results from a balance of opposing
effects of shortening of epicardial counterclockwise
fibers and endocardial clockwise fibers, with epicardial
fibers dominating. Relative endocardial ischemia due
to inotropic stimulation might lead to alteration in this
endocardial-epicardial force balance, and increased torsion. Similarly, if inotropic stimulation were to result in
a larger absolute force increment in the epicardial
compared with endocardial fibers, the balance would
shift and torsion would increase. The mechanisms of
this force-dependent increase in torsion require further
study.
perfused, ejecting heart model in dogs, enabling study
of the individual effects of preload, afterload, and
contractility on torsion while maintaining other factors
constant. Changes in load (preload or afterload) result
in changes in systolic torsion, which are mediated
through alteration in SV and EF. Dobutamine infusion
increases systolic torsion associated with an increase in
PSP even when volumes are identical to those at
control, indicating that there is an inotropic effect on
cardiac torsion that is independent of, rather than
mediated through, changes in volume, but is force
dependent. These relationships may have clinical implications. We speculate that the torsion-volume relationship might provide a load-independent measure of
contractility that could be measured noninvasively.
This potential application requires further study.
Limitations
We thank Stephanie Bosley for image acquisition and Kenneth
Rent for surgical assistance.
This study was supported by National Heart, Lung, and Blood
Institute Grants RO1-HL-46223 (E. P. Shapiro) and RO1-HL-43722
(J. L. Weiss). Sheng-Jing Dong was a recipient of a research fellowship of Heart and Stroke Foundation of Canada.
Address for reprint requests and other correspondence: E. P.
Shapiro, Div. of Cardiology, Johns Hopkins Univ. School of Medicine,
Johns Hopkins Bayview Medical Center, 4940 Eastern Ave., Baltimore, MD 21224 (E-mail: [email protected]).
This work was presented in part at the 69th Annual Scientific
Session of the American Heart Association, New Orleans, LA, November 1996.
The combined use of MRI tagging and the isolated
heart model is unique. Torsion can be measured nondestructively, unlike with implanted beads, and factors
such as preload, afterload, and contractility can be
controlled independently, compared with the model
with intact circulation. However, there exist some
limitations. Currently, MRI acquisition requires a
steady state, and therefore rapid beat-by-beat variation
of torsional dynamics during an intervention cannot be
measured. Secondly, the isolated heart model has some
nonphysiological aspects. Most importantly, the heart
rates of these isolated hearts were high; the presence of
incomplete relaxation or rate-related changes in contractility might make the results differ from those observed
at lower heart rates. Also, the mitral ring is fixed and
the chordae tendineae are severed; DeAnda et al. (8)
reported reduced torsion in dogs with mitral valve
replacement. However, when expressed in similar units,
the magnitude of torsion under baseline conditions of
the present study is quite similar to those reported by
Buchalter et al. (6) and Beyar et al. (5) in intact dogs
and is greater than that reported by DeAnda et al. (8) in
dogs with mitral valve replacement, suggesting that
the extent of twist was not greatly disturbed in this
model. Third, in this experiment we varied a computer
simulation of preload and afterload, which changed
end-diastolic and end-systolic volumes. Our model only
approximates the true afterload seen by the LV myocardium because it does not consider longitudinal and
circumferential radius of curvature, wall thickness, etc.
However, the secondary volume changes were precisely
controllable and easily measurable and therefore the
volumes (in ml) rather than loads are reported here.
Fourth, we attribute the inotropic-related change in
torsion at constant volume to differences in force development; we cannot exclude the possibility that other
effects of dobutamine that were not measured here,
such as velocity of shortening and ejection, may also
have contributed. Finally, although it has been demonstrated that there is a regional variation in torsion (6, 7,
13, 14, 20), we used global values in this study.
In conclusion, MRI tagging allowed us to measure
systolic torsion nondestructively in the isolated, blood-
Received 15 June 1998; accepted in final form 31 May 1999.
REFERENCES
1. Akima, H. A new method of interpolation and smooth curve
fitting based on local procedures. J Assoc. Comp. Machinery 17:
589–602, 1970.
2. Arts, T., W. C. Hunter, A. S. Douglas, A. M. Muijtjens, J. W.
Corsel, and R. S. Reneman. Macroscopic three-dimensional
motion patterns of the left ventricle. Adv. Exp. Med. Biol. 346:
383–392, 1993.
3. Arts, T., S. Meerbaum, R. S. Reneman, and E. Corday.
Torsion of the left ventricle during the ejection phase in the intact
dog. Cardiovasc. Res. 18: 183–193, 1984.
4. Arts, T., and R. S. Reneman. Measurement of deformation of
canine epicardium in vivo during cardiac cycle. Am. J. Physiol.
239 (Heart Circ. Physiol. 8): H432–H437, 1980.
5. Beyar, R., F. C. P. Yin, M. Hausknetch, M. L. Weisfeldt, and
D. A. Kass. Dependence of left ventricular twist-radial shortening relations on cardiac cycle phase. Am. J. Physiol. 257 (Heart
Circ. Physiol. 26): H1119–H1126, 1989.
6. Buchalter, M. B., F. E. Rademakers, J. L. Weiss, W. J.
Rogers, M. L. Weisfeldt, and E. P. Shapiro. Rotational
deformation of the canine left ventricle measured by magnetic
resonance tagging: effects of catecholamines, ischaemia, and
pacing. Cardiovasc. Res. 28: 629–635, 1994.
7. Buchalter, M. B., J. L. Weiss, W. J. Rogers, E. A. Zerhouni,
M. L. Weisfeldt, R. Beyar, and E. P. Shapiro. Noninvasive
quantification of left ventricular rotational deformation in normal humans using magnetic resonance imaging myocardial
tagging. Circulation 81: 1236–1244, 1990.
8. DeAnda, A., Jr., M. Komeda, S. D. Nikolic, G. T. Daughters
II, N. B. Ingels, Jr., and D. C. Miller. Left ventricular function,
twist, and recoil after mitral valve replacement. Circulation 92:
II-458–II-466, 1995.
9. Dong, S. J., S. A. Buffer, Jr., P. S. Hees, K. Rent, J. L. Weiss,
and E. P. Shapiro. Time-dependent interactions between epicardial and endocardial fibers determine left ventricular torsion
(Abstract). J. Am. Coll. Cardiol. 27, Suppl.: 32A, 1996.
10. Gibbons Kroeker, C. A., H. E. D. J. ter Keurs Hedj, M. L.
Knudtson, J. V. Tyberg, and R. Beyar. An optical device to
H1060
11.
12.
13.
14.
15.
16.
17.
18.
19.
20.
DETERMINANTS OF LEFT VENTRICULAR TORSION
measure the dynamics of apex rotation of the left ventricle.
Am. J. Physiol. 265 (Heart Circ. Physiol. 34): H1444–H1449,
1993.
Gibbons Kroeker, C. A., J. V. Tyberg, and R. Beyar. Effects of
load manipulations, heart rate, and contractility on left ventricular apical rotation. An experimental study in anesthetized dogs.
Circulation 92: 130–141, 1995.
Halperin, H., J. Tsitlik, R. Beyar, N. Chandra, and A.
Guerci. Intrathoracic pressure fluctuations move blood during
CPR: comparison of hemodynamic data with predictions from a
mathematical model. Ann. Biomed. Eng. 15: 385–403, 1987.
Hansen, D. E., G. T. Daughters II, E. L. Alderman, N. B.
Ingels, E. B. Stinson, and D. C. Miller. Effect of volume
loading, pressure loading, and inotropic stimulation on left
ventricular torsion in humans. Circulation 83: 1315–1326, 1991.
Hansen, D. E., G. T. Daughters II, E. L. Alderman, E. B.
Stinson, J. C. Baldwin, and D. C. Miller. Effect of acute
human cardiac allograft rejection on left ventricular torsion and
diastolic recoil measured by intramyocardial markers. Circulation 76: 998–1008, 1987.
Ingels, N. B., D. E. Hansen, G. T. Daughter II, E. B. Stinson,
E. L. Alderman, and D. C. Miller. Relation between longitudinal, circumferential, and oblique shortening and torsional deformation in the left ventricle of the transplanted human heart.
Circ. Res. 64: 915–927, 1989.
Judd, R. M., and B. I. Levy Effects of barium-induced cardiac
contraction on large and small vessel intramyocardial blood
volume. Circ. Res. 68: 217–225, 1991.
Kass, D. A., and W. L. Maughan. From Emax to pressurevolume relations: a broader view. Circulation 77: 1203–1212,
1988.
Katz, A. M. Influence of altered inotropy and lusitropy on
ventricular pressure-volume loop. J. Am. Coll. Cardiol. 11:
438–445, 1988.
LeGrice, I. J., Y. Takayama, and J. W. Covell. Transverse
shear along myocardial cleavage planes provides a mechanism
for normal systolic wall thickening. Circ. Res. 77: 182–193, 1995.
MacGowan, G. A., D. Burkhoff, W. J. Rogers, D. Salvador,
H. Azhari, P. S. Hees, J. L. Zweier, H. R. Halperin, C. O. Siu,
21.
22.
23.
24.
25.
26.
27.
28.
29.
J. A. C. Lima, J. L. Weiss, and E. P. Shapiro. Effects of
afterload on regional left ventricular torsion. Cardiovasc. Res.
31: 917–925, 1996.
Moon, M. R., N. B. Ingels, Jr., G. T. Daughters II, E. B.
Stinson, D. E. Hansen, and D. C. Miller. Alterations in left
ventricular twist mechanics with inotropic stimulation and
volume loading in human subjects. Circulation 89: 142–150,
1994.
Olsen, C. O., J. S. Rankin, C. E. Arentzen, W. S. Ring, P. A.
McHale, and R. W. Anderson. The deformational characteristics of the left ventricle in the conscious dog. Circ. Res. 49:
843–855, 1981.
Rademakers, F. E., M. B. Buchalter, W. J. Rogers, E. A.
Zerhouni, M. L. Weisfeldt, J. L. Weiss, and E. P. Shapiro.
Dissociation between left ventricular untwist and filling. Accentuation by catecholamines. Circulation 85: 1572–1581, 1992.
Rogers, W. J., J. L. Zweier, W. H. Guier, H. Azhari, W. L.
Graves, E. P. Shapiro, and J. L. Weiss. High resolution
analysis of cardiac deformation using an internal loop gap
resonator (Abstract). SMRM Book of Abstracts. 2: 861, 1991.
Suga, H., and K. Sagawa. Instantaneous pressure-volume
relationships and their ratio in the excised, supported canine left
ventricle. Circ. Res. 35: 117–126, 1974.
Sunagawa, K., D. Burkhoff, K. O. Lim, and K. Sagawa.
Impedance loading servo pump system for excised canine ventricle. Am. J. Physiol. 243 (Heart Circ. Physiol. 12): H346–H350,
1982.
Tsitlik, J., H. Halperin, S. Popel, A. Shoukas, F. C. P. Yin,
and N. Westerhof. Modeling the circulation with threeterminal electrical networks containing special nonlinear capacitors. Ann. Biomed. Eng. 20: 595–616, 1992.
Young, A. A., C. M. Kramer, V. A. Ferrari, L. Axel, and N.
Reichek. Three-dimensional left ventricular deformation in
hypertrophic cardiomyopathy. Circulation 90: 854–867, 1994.
Zerhouni, E. A., D. M. Parish, W. J. Rogers, A. Yang, and
E. P. Shapiro. Human heart: tagging with MR imaging. A
method for non-invasive assessment of myocardial motion. Radiology 169: 59–63, 1988.